Conventional designs of spinal interbody fusion cages have mainly focused on providing immediate strength to maintain disc height and shielding bone grafts within the cage. As such, the geometric features of conventional designs show little distinction from each other and most designs fall into a category consisting of pipe shapes with thick shells as outer walls and a hollow interior space that brackets the fill of grafting materials.
For example, the following interbody fusion devices were tested in a study conducted by Kanayama: (A) the BAK device, a titanium threaded cage (D=13 mm, L=20 mm); (B) the BAK Proximity device, a titanium threaded cage (D=13 mm, L=20 mm); (C) the RAY TFC device, a titanium threaded cage (D=14 mm, L=21 mm); (D) the Danek TIBFD device, a stainless steel threaded cage (D=16 mm, L=26 mm); (E) the single oval Harms device, a titanium cylindrical mesh cage (17 mm×22 mm×13 mm); (F) the double oval Harms device, a titanium cylindrical mesh cage (D=14 mm, L=13 mm); (G) the Brantigan PLIF device, a carbon fiber rectangular cage (13 mm×13 mm×24 mm); (H) the Brantigan ALIF device, a carbon cylindrical rectangular cage (24 mm×35 mm×13 mm); (I) a femoral ring allograft device, a sliced femoral shaft (20 mm×24 mm×14 mm); (J) a bone dowel, a dowel-shaped allograft with one hole (D=14 mm, L=18 mm); and (K) the In Fix device, a titanium cylindrical implant (20 mm 29 mm×15 mm).
Conventional approaches can be further divided into sub-groups defined by the threaded or non-threaded anchoring mechanism that the cage devices rely on to form rigid bonds with vertebral bodies. Threads may be provided along the entire outer surface of cylindrical cages, or may only be provided on two sides of wedge shaped cages.
Conventional hollow pipe designs guarantee sufficient reconstruction stiffness in arthrodesis and play a substantial role in stability for motion segments postoperatively. Nonetheless, the rigid shells may shield an implanted graft or ingrown bone tissue from sufficient mechanical stimulus, (known as “stress-shielding”) thus increasing the risk for decreased mineralization and bone resorption and provide a stress-shielded environment inside the device. The concern for stress shielding has been widely discussed and investigated. The decreasing bone mineral density is believed to be attributable to the lack of mechanical stimuli which increases the risk of bone resorption.
In view of the foregoing, many modifications have been made to reduce the effect of stress shielding. A common approach was to adjust the pore size and the pore distribution on the shielding shell. It was concluded that the stress shielding effect was correlated with the largest pore size rather than the total porous area. However, it is important to note that increasing the largest pore size sacrifices the stiffness of the cage. This may yield excess compliance causing unexpected deformation and instability.
In general, the requirement for a shell thickness sufficient to be capable of carrying spinal loads leads to stress shielding. Conventional designs do not have the flexibility to meet the multiple design requirements necessary to achieve sufficient rigidity, reduced stress shielding, and large porosity for biofactor delivery.
Another approach which increases flexibility is by changing the base material to carbon fiber or a cortical bone allograft shaft. However, the material substitutes bring up additional issues of biocompatibility and immune response. In addition, there may be a limited supply of cortical bone allograft.